Implantable medical devices such as pacemakers, defibrillators, and implantable cardioverter defibrillators (“ICDs”) have been successfully implanted in patients for years for treatment of heart rhythm conditions. Pacemakers are implanted to detect periods of bradycardia and deliver low energy electrical stimuli to increase the heart rate. ICDs are implanted in patients to cardiovert or defibrillate the heart by delivering high energy electrical stimuli to slow or reset the heart rate in the event a ventricular tachycardia (VT) or ventricular fibrillation (VF) is detected. Another type of implantable device detects an atrial fibrillation (AF) episode and delivers electrical stimuli to the atria to restore electrical coordination between the upper and lower chambers of the heart. Still another type of implantable device stores and delivers drug and/or gene therapies to treat a variety of conditions, including cardiac arrhythmias. The current generation for all of these implantable devices are typically can-shaped devices implanted under the skin that deliver therapy via leads that implanted in the heart via the patient's vascular system.
Next generation implantable medical devices may take the form of elongated intravascular devices that are implanted within the patient's vascular system, instead of under the skin. Examples of these intravascular implantable devices are described, for example, in U.S. Pat. No. 7,082,336, U.S. Publ. Appl. Nos. 2005/0043765A1, 2005/0208471A1 and 2006/0217779A1. Devices of this type can have a diameter of about 3-15 mm and a length of about 10-60 cm to facilitate insertion and implantation inside of the vasculature, while permitting a sufficient amount of blood flow around the device. Within geometric constraints such as these, the devices contain electrical/electronic components and circuitry for performing their various functions.
Implantable devices have on-board energy storage (typically, batteries), and high voltage converter circuits for converting the stored energy into a form suitable for operating the device to deliver electrotherapy therapy. In cardioverter/defibrillator-type devices, the high voltage converter circuitry typically includes a circuit that produces energy at high voltage (typically at least in the range of 50-800 Volts and 1-40 Joules) for use in the application of the cardioversion/defibrillation electrotherapy. Because there is only a finite amount of energy available in the energy storage, and because replacing the batteries typically involves a surgical procedure to remove or otherwise access the implanted device or a recharge process that can require extended periods of time for recharging the energy storage, providing highly efficient circuitry is important to prolonging the useful life of the device and also to making the device as small as practicable. Accordingly, the high voltage converter circuit used in implantable devices should be as efficient as possible.
A switching mode power converter is generally considered to be one of the most efficient arrangements for stepping up voltage from the energy storage to the high voltage required for delivery of the electrotherapy. This type of converter operates by applying intermittent current to an inductive element such as a choke or a transformer, and harnessing the voltage-boosting effect produced by the associated time-varying magnetic field generated by the inductive element. A variety of switching converter topologies and operating modes are well-known. Examples include the boost converter, the flyback converter, the SEPIC (single-ended primary inductance converter), and the Cuk converter. The boost converter and certain Cuk converter topologies use one or more inductors, whereas the flyback, SEPIC, and other types of Cuk converters use transformers as the principal inductive elements for performing the voltage conversion function. Certain SEPIC topologies use both, an inductor, and a transformer.
The inductive element (whether an inductor or a magnetically coupled set of inductors) is generally constructed from at least one coil of wire and a magnetic core of high relative permeability material, such as ferromagnetic material. The core operates to confine the magnetic field closely to the element, thereby increasing its inductance. The core provides a magnetic flux path that guides the flux through the center of the coil(s) and along a return path that can be contiguous, or can alternatively have a plurality of non-contiguous return path portions. A variety of core geometries are known for inductive elements. Some are constructed as enamel coated wire wrapped around a ferrite bobbin with wire exposed on the outside, while others enclose the wire completely in ferrite for improved shielding effect. Core geometries typically include toroidal structures, C- or E-shaped structures, pot-shaped structures and planar structures.
In the case of a switching mode transformer, a typical turns ratio for use in a high voltage converter circuit for an implantable device can be on the order of Np:Ns being 1:15, where Np is the number of primary turns and Ns is the number of secondary turns. Unlike transformers used for signals and linear power supplies, transformers used in switching mode circuits are designed not only to transfer energy, but also to store the energy for a significant fraction of the switching period. For instance, in a power converter switching at about 60 kHz (which is a frequency selected to keep core eddy current losses low) and having a transformer with a core made from a power ferrite material with relative permeability of 2000 to 4000, a certain minimum primary inductance is required in the transformer.
Most of the stored energy in an inductive element is stored in an air gap of the core. A certain air gap volume is needed to store the desired energy. However, increasing the gap length reduces the inductance in the transformer or inductor. Winding inductance in an inductive element is directly proportional to the square of the number of windings, and to the magnetic cross sectional area orthogonal to the direction of magnetic flux produced in the volume. To compensate for the loss of inductance due to an increased air gap, a greater number of windings or a greater cross-sectional area for the magnetic flux path is needed. More windings take up more volume, and increase the power losses in the device due to increased resistance. Increasing the cross-sectional area for the magnetic flux path in a conventional core geometry would involve increasing the size of the core and consequently taking space away from the windings or increasing the overall size of the device.
In terms of an intravascular implantable device which may take the form of an elongated structure implanted within a patient's vasculature and generally having a circular cross-sectional area, if a standard circular pot core is used as the ferrite core of the transformer, the magnetic cross sectional area will be limited to something less than the cross-sectional area of the implantable device. Given this limitation, one alternative to increasing inductance is to increase the number of windings. Unfortunately, this adds to the winding volume in the transformer as a relatively high windings turns-ratio is needed for the high voltage converter. Aside from the higher overall resistance in the windings by increasing the total number of turns, this approach would also require a longer transformer to accommodate the windings.
A long and narrow pot core poses difficult winding challenges when used in an implantable intravascular device owing to the limited winding cross sectional area across the diameter of the core. Furthermore, there is a practical limit to the length of the transformer in implantable intravascular devices. For instance, the housing of the implantable intravascular device must provide a certain amount of flexibility to facilitate routing of the device through the vasculature. Longer sections of rigid housing elements limit the flexing radius of the device. In addition, the enclosure section housing the transformer may need space beyond the ends of the transformer to house circuitry, input/output hardware, wiring, and the like.
Other approaches, such as scaling down an E core or one of its derivatives, such as the EFD or ER cores, for use within the dimensional confines of an intravascular device may not be feasible given the energy storage and inductance requirements for the power converter circuit. For instance, there may be insufficient winding area to achieve the target primary inductance for a transformer. Even if the electrical performance were achievable in the small size, using a scaled-down E core-type inductive element in the intravascular device's housing would be wasteful of housing volume because excess volume would remain in the housing around the inductive element.
Given the size constraints of intravascular implantable devices, designing a power converter that can effectively and efficiently generate the high voltage electrotherapy signals using present-day inductive elements presents significant challenges. Typical core shapes and geometries, such as the E, C, toroidal, and pot cores ordinarily capable of providing the required functional and performance requirements for high voltage converters in conventional implantable device like conventional can-shaped implantable defibrillators are not well-suited for use in the small-diameter space of implantable intravascular devices.